Artificial or prosthetic devices for replacing defective joints in humans, particularly the hip joint, have been the subject of extensive research and development efforts for many years. In total hip arthroplasty, the most common adult reconstructive hip procedure currently performed in the U.S., a metallic femoral component is typically inserted into the natural medullary cavity of the femur. Simultaneously, an acetabular cup, usually of high-density polyethylene, is inserted into the acetabulum.
A typical prior art femoral component 2 is shown in FIG. 1. The component 2 is an integral metallic component having a head 4, a neck 6 and a stem 10 having a medial side 11 and a lateral side 13. There is usually a collar 7 between the neck 6 and the stem 10. The medial extension of the collar 7 is the platform 8. The stem 10 has a proximal end 19 and a distal end 20 which ends at the tip 22. Various means of measuring such femoral components are used. The neck length 12 is measured from the center 18 of the head 4 to the base of the collar 7. The head-stem offset 14 is measured from the center 18 of the head 4 to the line 22 through the axis of the distal part 20 of the stem 10. The stem length 16 is measured from the medial base of the collar 7 to the tip 22 of the stem 10. The angle .alpha. of the neck 6 is measured by the angle at the intersection of the line 24 through the center 8 of the head 4 and the neck 6 with another line 26 along the lateral border of the distal half 20 of the stem 10.
The femoral component may be made of any strong inert material. Materials which have been used in the past on such components include stainless steel, chromium cobalt molybdenum alloy (Co-Cr-Mo), titanium, or a combination such as Co-Cr-Mo with a ceramic head or titanium with a cobalt-chromium or ceramic head. It may also be made of isoelastic polyacetate.
The head diameter is usually either 22, 26, 28, 32, or 38 mm with a neck length of 30-42 mm. The cross-section of the neck may be round, oval, or trapezoidal. The collar itself may or may not be present. The surface of the stem may be polished, dull, pre-coated with cement, press-fit, or have a porous-metal coating. There may or may not be fenestrations in the stem. The proximal third of the stem may be curved or angulated. The stem may be sabre-shaped, tapered, have a straight lateral edge or an anterior bow or a wide proximal third. The head-stem offset is generally 38-45 mm and the length is generally 12-18 cm or longer. Sometimes the femoral component is made as a modular system with a tapered metal post on the stem to mate with a head component that makes for different neck lengths and diameter of heads made of cobalt-chrome or ceramic. Reference is made to Calandruccio, R. A., "Arthroplasty of Hip"in Campbell's Operative Orthopaedics, Vol. 2, St. Louis, C. B. Mosby, 1987, chapter 41, pages 1213-1501.
A major problem from which most prior art femoral components suffer is stability of the component in place. Lack of complete stability can cause pain, failure of the artificial hip, fracture of the femur, or various other problems. Many attempts have been made to avoid such problems and add stability. One such attempt is the use of grouting medium or bone cement to fix the femoral component to the bone. In this case, bone is cleared from the medullary cavity to produce a larger space than required for the stem. Grouting material is inserted to fill the gap between the bone and the stem, as a means for fixing the device and as a means for load transfer between the device and the remaining bone.
While such a method is advantageous in that accurate insertion into the bone is not required and immediate mechanical fixation can be achieved leading to early weight bearing and rehabilitation, many disadvantages result from the inherent weakness of the cement which is exacerbated by poor distribution and/or contamination by blood during surgery.
Efforts have been made to fix implants without the use of a grouting medium, in which case it is important that an accurate bone resection be performed. The femoral component must be selected to give the tightest fit possible to provide a mechanically stable support for physiological loading.
Sometimes the surface of the implant is treated to provide a porous or roughened structure which acts to promote bone tissue growth around the implant, further stabilizing the femoral component with respect to the bone.
A major advantage of the latter system is the absence of cement or grouting medium, thus eliminating the long term inherent weakness and the short term toxic effects of these materials. The disadvantages are numerous. First, these stems have the added requirement of a sufficiently tight fit to prevent motion between metal and bone. Accurate bone resection customized to each type of available implant is difficult to achieve and often results in some initial looseness or lack of support. The implant will subsequently migrate to a more stable position, which may not be the ideal orientation for proper function of the femoral component. The requirement for a tight fit increases the possibility of fracture of the femur during insertion. Additionally, the patient must avoid bearing full weight on the hip for approximately six weeks to allow for bone formation.
Treatment of the implant to form the porous or roughened surfaces may cause local stress sites in the implant which significantly increase the risk of fatigue fractures. Further, a considerable time is required for bone tissue ingrowth and stabilization of the implant to occur. This is a significant detriment to early patient rehabilitation. Additionally, surface treatment exposes a greater surface area of the implant, increasing diffusion of metal ions which are associated with an increased risk of toxic or pathological effects.
The implant's stem may weaken from improper stress loads or decreased fatigue strength due to surface treatment. If this happens, the stem may bend or fracture, requiring its removal, which is particularly difficult if significant bone growth has occurred.
Various efforts have been made to design a femoral component hip endoprosthesis that can be implanted in the medullary canal in such a way as to provide implant stability without resorting to surface treatment. Some such efforts are directed to creating an isoelastic femoral component which is adaptable to the shape of the cavity created for the prosthesis in the femur and thus transfers the load from the implant outward to the bone surrounding the femoral component in the medullary canal. See, for example, U.S. Pat. No. 4,743,263. Other implants use stepped projections U.S. Pat. No. 4,031,571). Or fixation wires (U.S. Pat. No. 4,530,114) to impose tensile forces on the lateral side of the femoral component in the medullary canal. This is reportedly done to anchor the femoral component while taking advantage of the natural conditions of the bone.
Other efforts to stabilize implants have been directed to adding pins or studs (U.S. Pat. No. 3,896,505), wing-like extensions to prevent rotation of the shank (U.S. Pat. No. 4,664,668), plates to provide anti-rotation stability for the implants (U.S. Pat. No. 4,904,269) and anti-rotation fins (U.S. Pat. No. 4,936,863). All of these efforts are directed to preventing the femoral component from rotating inside the medullary canal after insertion as force is applied to the implant by the patient returning to his or her feet.
Most of these implants suffer the disadvantage of pros thesis dislocation and bone fracturing due to improper force distribution on the femur.
Prior to the present invention, all implants have been designed based on the conventional assumption that the lateral femur is under tensile stress when unilateral loading forces are applied to the femur head. This assumption is based on the standard model for describing the biomechanics of the human hip described in the classic work of Koch, published in 1917 (Am. J. Anat. 21:177, 1917). He determined that the medial aspects of the femur are under compressive load during unilateral load, such as during a stride, and most of the lateral cortex is under tensile loading. In Koch's model, most of the force generated within the hip is attributed to the load of the abductor muscles, anatomically defined as taking origin from the lateral aspect of the iliac crest of the hip bone and inserting on the greater trochanter of the femur. Thus, the superimposed body weight creates a lever across the head of the femur, which serves as a fulcrum, with the body weight force being balanced by the abductor muscle force. This model leads to the conventional wisdom that the lateral aspects of the upper femur are under tensile loading.
In the design of femur components, it is advisable to subject every stem of each new design to static and dynamic testing. Such testing is necessary to ensure that a particular design does not fail prematurely due to fatigue. Thus, a need exists within the medical equipment industry to assess the endurance properties of femoral components of hip replacements and to provide a standard against which they can be prepared. The standards which have been used to date, however, are all based on the Koch model.
The present standards and proposed standards are described in Humphreys, P. K. et al., "Testing of Total Hip Replacements: Endurance Tests and Stress Measurements Part I: Endurance Tests", Proc. Instn. Mech. Engrs., 204:29-34 (1990) and Humphreys, P. K., et al., "Testing of Total Hip Replacements: Endurance Tests and Stress Measurements Part II: Stress Measurements, Proc. Instn. Mich. Engrs., 204:35-41 (1990). Such measuring systems are also disclosed in Semlitsch, M. et al., "Ten Years of Experience With Test Criteria for Fracture-Proof Anchorage Stems of Artificial Hip Joints", Engineering and Medicine, 12:185-198 (1983). A diagram showing this conventional stress measurement system is shown in FIG. 15. In accordance with this testing procedure, the head 110 of the femoral component 112 to be tested is firmly clamped and inserted into a specimen holder 114 aligned in accordance with predetermined orientation angles. A fixing medium 116 is then poured into the holder 114 until it reaches a predetermined depth, usually about 50 mm below the collar of the stem. The specimen 112 is then left while the embedding medium 116 hardens. Epoxy resin is usually used as the embedding medium. The holder 114 with the fixed specimen 112 is then located in the testing machine and the entire stem may be immersed in a saline bath (not shown) in order to maintain physiological temperatures. The testing machine 119 then applies a load 118 vertically downward upon the head 110 of the femoral component of a predetermined amount and with a predetermined frequency. The vertical deflection is measured during the first minute of testing. A computer program controlling the load cycle is adjusted so that if the deflection exceeds 120% of the initial deflection, the testing machine will stop. Strain gauges 120 may be placed along the medial and lateral surfaces of the component stem 122 and the maximum bending stresses measured.
Dobbs, H. S. "A Model Femur for In Vitro Testing of Femoral Components", J. Biomed. Eng., 3:225-34 (1981) discloses another model femur system for testing femoral components. This device uses a thin-walled tubular fixture of appropriate dimensions to simulate the femur. Strain gauges are placed on the stem as well as on the tubular structure. A three point bending test stand is disclosed in Reubin, J. D. et al., "Comparative Mechanical Properties of Forty-Five Total Hip Systems", Clin. Orthop., 141:55-65 (1979).
In Tanner, K. E. et al., "A System for Modeling Forces on the Hip Joint in One-Legged Stance", J. Biomed. Eng., 10:289-90 (1988), the authors recognize that the previous systems for testing of hips or hip prostheses using a single unidirectional load do not make allowances for the loads applied via the greater or lesser trochanters. Accordingly, Tanner proposed an experimental arrangement which more closely models the forces on the hip joint in one-legged stance, taking into account the abductor muscles between the iliac crest and the greater trochanter. This test system is shown in FIG. 16. This model uses a 200 mm U channel 126 to model the pelvis with a movable plastic "acetabulum" 128. A proximal section of a femur 130, suitably reamed, is implanted with an uncemented femoral component 132 and mounted vertically in a cylindrical holder 134, 150 mm from the line of action of the applied load. The abductors were modeled by a strip of stainless steel braid 136, screwed to the lateral aspect of the greater trochanter 140 using long cancellous bone screws 138. The braid 136 was then held at the end 142 of the U channel 126 at 20.degree. to the vertical. Thus, when a load 144 was applied to the U channel 126 modeling the pelvis, the channel 126 remained approximately horizontal. The acetabular cup 128 was positioned to ensure that the femoral head 146 fitted into it with the "abductor tendon" 136 at the appropriate angle. Thus, when 83.5% of body weight was applied via the load cell 148 of a Schenk Trebel testing machine, the forces in the femur 130 and in the prosthesis 132 were supposed to be equivalent to those applied by standing on one leg so that the compressive force through the femoral head is reacted by a tensile force from the greater trochanter.